The main MRI applications are in the field of medical imaging. It is important to provide the best RF coil performance—the highest RF magnetic field per unit power delivered to the coil, while simultaneously obtaining the highest RF field homogeneity—in order to obtain fast and reliable MR images of a patient or volunteer subject.
For magnetic resonance imaging (MRI), especially at high magnetic field level, it has become important to use a multi-channel transmit coil for producing the desired RF field. The coil can be excited by a single channel transmit power source followed by a multi-channel power splitter, or by a multi-channel power transmitter unit. Due to the natural geometrical electromagnetic (EM) interaction, it is impossible to design a multi-channel coil without EM coupling of its elements. In a first approximation this EM coupling can be characterized by the S parameter coefficients Sxy where x and y correspond to coil element input numbers.
Several EM coupling compensation methods have been proposed to address the problem.
One category of methods introduces partial geometric overlap of coil elements to annul the mutual inductance between them (cf. ROEMER P B, EDELSTEIN W A, HAYES C E, SOUZA S P, MUELLER O M.: “The NMR phased array”, Magn Reson Med 1990, 16:192-225; ALAGAPPAN V, WIGGINS G, POTTHAST A, SETSOMPOP K, ADALSTEINSSON E, WALD L.: “An 8 channel transmit coil for transmit SENSE at 3 T”, Proceedings of the 14th Annual Meeting of ISMRM, Seattle, Wash., USA, 2006). Such methods are effective for nearest-neighbor elements only, and tend to impose stringent constraints on the geometry and placement of the individual coils.
Another category of methods employs a capacitive, inductive or a multi-port decoupling network at the cost of increased RF loss and increased complexity of the decoupling circuits and tuning efforts (cf. “The NMR phased array”, l.c.; “An 8 channel transmit coil for transmit SENSE at 3 T”, l.c.; WANG J.: “A novel method to reduce the signal coupling of surface coils for MRI. In: Proceedings of the 4th Annual Meeting of ISMRM, New York, N.Y., USA, 1996).
A third category of methods suppresses the coupling-induced currents with high source impedance, for example, driving non-resonant loop-shaped coils directly by power metallic oxide semiconductor field effect transistors (cf. “The NMR phased array”, l.c.). However this technology has not been implemented in commercial available MRI scanner yet.
The major challenge in multi-channel coil design consists in the difficulty (or in most cases impossibility) of designing a solution in which the EM coupling compensation performance is independent of coil loading. In addition, for relatively high frequency coils where the RF wavelength in the load is comparable or less than the coil element length, the electromagnetic coupling becomes spatially distributed. As a result, such coupling (defined here as distributed EM coupling, as opposed to lumped EM coupling as found in low frequency coils) cannot be described using only lumped element theory, it cannot be estimated explicitly as a value for Sxy, and it cannot be fully compensated using approaches derived from lumped element based analysis.
By considering power balance, it is clear that a coil's magnetic field approaches a maximum when the RF power reflected by the entire coil (PRef, Coil) approaches its minimum at the MRI resonance frequency (fres) (cf. “The NMR phased array”, l.c.). However, there is no standardized commercial device for measuring PRef, Coil of multi-channel near-field RF coils that can be used for its minimization. The most common RF coil tuning procedure is the minimization of each element's reflection coefficient (Sxx). This usually entails that the Sxx of each element approaches its minimum at a frequency (fmin, Sxx) equal to fres. Sxx is measured by direct connection of a vector network analyzer (VNA) to each coil element's input while the other elements are terminated simultaneously by 50 Ohm loads. This works reliably when the difference of the frequencies at which Sxx and PRef, Coil approach their minimum is significantly smaller than the ratio of fres to loaded coil quality factor (Qcoil). It is valid for coils with a moderate number of elements, lumped EM coupling and when both Sxx and Sxy are lower than −20 dB. The higher the Sxy value and the larger the number of coil elements, the larger the difference in frequency at which Sxx and PRef, Coil approach their minimum. This occurs because for a multi-channel coil PRef, Coil depends not only on the Sxx and Sxy magnitudes but also on the phase distribution both of Sxy and the power delivered to each coil element.
Another approach for optimization of multi-channel coil performance is to maximize the magnetic field generated by individual elements (cf. “The NMR phased array”, l.c.). However, as with the Sxx minimization strategy, the higher the Sxy value and the larger the number of coil elements, the larger the difference between the frequencies at which the magnetic field of an individual coil element approaches its maximum and PRef, Coil approaches its minimum.
For distributed EM coupling there is no simple criterion useful for estimation when the difference of the frequencies at which Sxx and PRef, Coil approach their minimum becomes considerable. As a result it is impossible to predict in advance that coil performance optimization guided by Sxx and Sxy will succeed.
Direct PRef, Coil minimization yields a simple method for achieving the maximum near-field magnetic field generated for unit delivered power. In general, PRef, Coil can be calculated by obtaining the full S parameter matrix of a coil (where the number of unknowns is equal to the square of the number of coil elements), together with knowledge of the power amplitude and phase delivered to each port input. Measuring of the full S parameter matrix for a multi-channel coil is a rather time consuming process. After each tuning/matching step, as well as for any movement of the coil load, the S parameter matrix has to be redefined. This makes such experimental coil optimization procedure extremely complex and unsuitable for real applications.
In additional, for coils using a power splitter, it is difficult to obtain the amplitude and phase distribution by means of an NVA measurement of the power delivered to each coil element by each splitter channel, because the phase distribution depends on coil input impedance, which is never exactly 50 Ohm, and in general varies with coil loading and tuning conditions.
It is not obligatory to compensate EM coupling completely during coil performance optimization. By adjusting the RF coil tuning and matching sub-circuits, a condition can be achieved where some RF coils are suitable for MRI purposes, although they lack distributed or lumped EM coupling compensation, or have residual uncompensated EM coupling despite use of a decoupling method. The performance of such coils can remain close to that achievable with the best reachable coupling compensation. We define EM coupling condition in such coils as weak and in other coil types as strong.
This is not merely an academic issue, since some commercial available multi channel coils are built and marketed without any coupling compensation at all (cf. “The NMR phased array”, l.c.).
Several methods have been developed to assist optimization of RF coil homogeneity. One of these is a so-called RF shimming procedure, which can be applied with multi-channel RF coils to improve RF field homogeneity in the object region to be investigated (region of interest, ROI). Adjustment of the amplitude and phase of transmit signals for each individual coil element eliminates the imperfections of near-field coil performance within an ROI that is in most cases much smaller than the entire coil load volume, simultaneously influencing PRef, Coil and coil performance, as a result. Because the maximum value of power delivered to each coil input is limited by the available hardware, it is impossible without limit to increase input power to compensate for degraded coil performance. Thus, for MRI techniques involving adjustment of the amplitude and phase of transmit signals, the coil tuning/matching condition must be optimized for best performance within the range of amplitudes and phases provided.
There are several approaches for measurement of power reflected from a single RF coil input. A commonly used method involves installing a RF coupler in front of the RF coil input and measuring coupler output signal by an oscilloscope or a power meter. Additionally, recent developments have made available a versatile transmit path monitor using an in-line vector RF current and voltage (I/V) sensor connected to a digital console (cf. “The NMR phased array”, l.c.).